U.S. patent number 8,311,182 [Application Number 12/887,584] was granted by the patent office on 2012-11-13 for system and method of notch filtration for dual energy ct.
This patent grant is currently assigned to General Electric Company. Invention is credited to Naveen Chandra, Bruno Kristiaan Bernard De Man.
United States Patent |
8,311,182 |
Chandra , et al. |
November 13, 2012 |
System and method of notch filtration for dual energy CT
Abstract
An imaging system includes an x-ray source that emits a beam of
x-rays toward an object to be imaged, a detector that receives the
x-rays attenuated by the object, a spectral notch filter positioned
between the x-ray source and the object, a data acquisition system
(DAS) operably connected to the detector, and a computer operably
connected to the DAS and programmed to acquire a first image
dataset at a first kVp, acquire a second image dataset at a second
kVp that is greater than the first kVp, and generate an image of
the object using the first image dataset and the second image
dataset.
Inventors: |
Chandra; Naveen (Kenosha,
WI), De Man; Bruno Kristiaan Bernard (Clifton Park, NY) |
Assignee: |
General Electric Company
(Schenectady, NY)
|
Family
ID: |
45817765 |
Appl.
No.: |
12/887,584 |
Filed: |
September 22, 2010 |
Prior Publication Data
|
|
|
|
Document
Identifier |
Publication Date |
|
US 20120069953 A1 |
Mar 22, 2012 |
|
Current U.S.
Class: |
378/5; 378/98.9;
382/131; 382/130; 378/98.12 |
Current CPC
Class: |
A61B
6/482 (20130101); A61B 6/03 (20130101); G21K
1/10 (20130101); A61B 6/4042 (20130101); A61B
6/4035 (20130101); A61B 6/06 (20130101); A61B
6/4435 (20130101) |
Current International
Class: |
A61B
6/00 (20060101); G06K 9/00 (20060101) |
Field of
Search: |
;378/4,5,98.9,98.12
;382/130 |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Taningco; Alexander H
Attorney, Agent or Firm: Ziolkowski Patent Solutions Group,
SC
Claims
What is claimed is:
1. An imaging system comprising: an x-ray source that emits a beam
of x-rays toward an object to be imaged; a detector that receives
the x-rays attenuated by the object; a spectral notch filter
positioned between the x-ray source and the object; a data
acquisition system (DAS) operably connected to the detector; and a
computer operably connected to the DAS and programmed to: acquire a
first image dataset at a first kVp; acquire a second image dataset
at a second kVp that is greater than the first kVp; and generate an
image of the object using the first image dataset and the second
image dataset.
2. The imaging system of claim 1 wherein the first kVp includes a
mean kVp that is less than a k-edge of the notch filter, and
wherein the second kVp includes a mean kVp that is greater than the
k-edge of the notch filter.
3. The imaging system of claim 1 comprising a bowtie filter
positioned between the x-ray source and the object, wherein the
bowtie filter includes the spectral notch filter.
4. The imaging system of claim 3 wherein the bowtie filter includes
a first bowtie at a first axial location and a second bowtie at a
second axial location, and wherein the spectral notch filter is
positioned in the bowtie at the first axial location.
5. The imaging system of claim 1 wherein the first kVp is
approximately 80 kVp and the second kVp is approximately 140
kVp.
6. The imaging system of claim 1 wherein the computer is programmed
to decompose the first image dataset and the second image dataset
into a first basis material image and a second basis material
image, wherein the first basis material image is one of an iodine
image and a water image.
7. The imaging system of claim 1 wherein the computer is further
programmed to: cause the x-ray source to emit a first beam of
x-rays at the first kVp; cause the x-ray source to emit a second
beam of x-rays at the second kVp; acquire the first image dataset
using the first beam of x-rays; and acquire the second image
dataset using the second beam of x-rays.
8. The imaging system of claim 1 wherein the spectral notch filter
includes a first k-edge material having a first k-edge and a second
k-edge material having a second k-edge, wherein the x-rays
attenuated by the object pass through both the first k-edge
material and the second k-edge material.
9. The imaging system of claim 1 wherein the spectral notch filter
includes a first k-edge material having a first k-edge and is
positioned between the x-ray source and the object during the
acquisition of the first image dataset at the first kVp and during
the acquisition of the second image dataset at the second kVp.
10. The imaging system of claim 1 wherein the computer is
programmed to generate the image by being programmed to convert
measured projections of the first image dataset and of the second
image dataset to density line-integral projections, form density
maps of basis materials based on the density line-integral
projections, and form a volume rendering of one of the basis
materials based on the density maps.
11. A method of dual energy CT imaging comprising: selecting a low
kVp potential and a high kVp potential for dual energy imaging;
selecting a k-edge filter based on the low kVp potential and the
high kVp potential and based on a k-edge of a material in the
k-edge filter; positioning the k-edge filter between a source and
an object to be imaged; and acquiring imaging data with the source
energized to the first kVp potential and with the source energized
to the second kVp potential.
12. The method of claim 11 wherein the step of selecting the k-edge
filter comprises selecting the k-edge filter such the k-edge of the
material in the filter is above a peak kVp of the low kVp potential
and below a peak kVp of the high kVp potential.
13. The method of claim 11 comprising positioning a bowtie filter
between the source and the object to be imaged, wherein the bowtie
filter includes the selected k-edge filter.
14. The method of claim 11 comprising: decomposing the imaging data
into a first basis material image and a second basis material
image; and generating a final image based on the first basis
material image and the second basis material image.
15. The method of claim 11 comprising: selecting a bowtie filter
having a first bowtie profile at a first axial location and a
second bowtie profile at a second axial location; wherein the
selected k-edge filter is within the bowtie filter at one of the
first axial locations and the second axial locations.
16. The method of claim 11 wherein the k-edge filter comprises a
first k-edge material and a second k-edge material, such that the
imaging data acquired with the source energized to the first kVp is
acquired from x-rays having passed through both the first k-edge
material and the second k-edge material.
17. The method of claim 11 wherein the step of acquiring the
imaging data comprises acquiring the imaging data while the k-edge
filter is positioned between the source and the object to be imaged
when the source is energized to the first kVp potential and when
the source is energized to the second kVp potential.
18. A method of dual energy CT imaging comprising: passing low kVp
x-rays through a k-edge notch filter to generate a first x-ray
spectrum; acquiring a first set of imaging data of an object using
the first x-ray spectrum; passing high kVp x-rays through the
k-edge notch filter to generate a second x-ray spectrum; acquiring
a second set of imaging data of the object using the second x-ray
spectrum; and generating an image using the first set of imaging
data and the second set of imaging data.
19. The method of claim 18 comprising selecting the k-edge notch
filter based on the low kVp and the high kVp.
20. The method of claim 18 comprising positioning a bowtie filter
between the object and an x-ray source that is used to generate the
low kVp x-rays and the high kVp x-rays, wherein the bowtie filter
includes the k-edge notch filter.
21. The method of claim 20 wherein the bowtie filter includes a
first bowtie profile, and a second bowtie profile that is axially
offset from the first bowtie profile, wherein one of the first and
second bowtie profiles includes the kedge notch filter.
22. The method of claim 18 wherein the step of generating the
images comprises generating the image by converting measured
projections of the first set of imaging data and of the second set
of imaging data to density line-integral projections, forming
density maps of basis materials based on the density line-integral
projections, and forming a volume rendering of one of the basis
materials based on the density maps.
Description
BACKGROUND
The present invention relates generally to diagnostic imaging and,
more particularly, to a system and method of basis material
decomposition having an increased separation of mean energies
between low and high kVp projections.
Medical imaging devices comprise x-ray systems, magnetic resonance
(MR) systems, ultrasound systems, computed tomography (CT) systems,
positron emission tomography (PET) systems, ultrasound, nuclear
medicine, and other types of imaging systems. Typically, in CT
imaging systems, an x-ray source emits a fan-shaped beam toward a
subject or object, such as a patient or a piece of luggage.
Hereinafter, the terms "subject" and "object" shall include
anything capable of being imaged. The beam, after being attenuated
by the subject, impinges upon an array of radiation detectors. The
intensity of the attenuated beam radiation received at the detector
array is typically dependent upon the attenuation of the x-ray beam
by the subject. Each detector element of the detector array
produces a separate electrical signal indicative of the attenuated
beam received by each detector element. The electrical signals are
transmitted to a data processing system for analysis which
ultimately produces an image.
Generally, the x-ray source and the detector array are rotated
about the gantry opening within an imaging plane and around the
subject. X-ray sources typically include x-ray tubes, which emit
the x-ray beam at a focal point. X-ray detectors typically include
a collimator for collimating x-ray beams received at the detector,
a scintillator for converting x-rays to light energy adjacent the
collimator, and photodiodes for receiving the light energy from the
adjacent scintillator and producing electrical signals
therefrom.
Typically, each scintillator of a scintillator array converts
x-rays to light energy. Each scintillator discharges light energy
to a photodiode adjacent thereto. Each photodiode detects the light
energy and generates a corresponding electrical signal. The outputs
of the photodiodes are then transmitted to the data processing
system for image reconstruction. Such typical systems, however, do
not include an ability to discriminate spectral energy content of
x-rays as they pass through an object being imaged.
However, as known in the art, dual or multi-energy spectral CT
systems have been developed that can reveal the densities of
different materials in an object and generate images acquired at
multiple monochromatic x-ray energy levels. In the absence of
object scatter, a system derives the behavior at a different energy
based on a signal from two regions of photon energy in the
spectrum: the low-energy and the high-energy portions of the
incident x-ray spectrum. In a given energy region of medical CT,
two physical processes dominate the x-ray attenuation: (1) Compton
scatter and the (2) photoelectric effect. The detected signals from
two energy regions provide sufficient information to resolve the
energy dependence of the material being imaged and the relative
composition of an object composed of two hypothetical
materials.
Different approaches have been developed to realize dual energy or
spectral imaging. To name a few, dual x-ray source and detector, a
single x-ray source with an energy discriminative detector, and a
single x-ray source and detector with multiple acquisitions at
different kVp or interleaved with fast kVp switching capability are
examples of techniques.
In a dual x-ray source and detector system, typically two x-ray
sources are provided, each having a respective detector positioned
opposite thereto such that x-rays may be emitted from each source
having a different spectral energy content. Thus, based on the
known energy difference of the sources, a scintillating or energy
integrating device may suffice to distinguish energy content and
different materials within the object being imaged.
In a single x-ray source with an energy discriminative detector,
energy sensitive detectors may be used such that each x-ray photon
reaching the detector is recorded with its photon energy. Such
systems may use a direct conversion detector material in lieu of a
scintillator.
In a single x-ray source and detector arrangement, a conventional
third generation CT system may acquire projections sequentially at
different peak kilovoltage (kVp) levels, which changes the peak and
spectrum of energy of the incident photons comprising the emitted
x-ray beams. Two scans are acquired--either (1) back-to-back
sequentially in time where the scans require two rotations around
the subject, or (2) interleaved as a function of the rotation angle
requiring one rotation around the subject, in which the tube
operates at, for instance, 80 kVp and 140 kVp potentials.
When dual energy data is acquired back-to-back, imaging data
acquired during subsequent source/detector gantry rotations is
prone to motion artifacts because of the motion that occurs during
each subsequent rotation. When interleaved, in contrast, an input
voltage to the x-ray source is switched quickly between the low and
high kVp potentials, which allows a close correlation between
imaging data sets. However, because the switching occurs very
rapidly on a single x-ray source, there is little opportunity to
change the filtration between the two samples. As a result, there
is a spectral (energy) overlap between the two samples that
inherently limits the amount of energy separation between them. As
known in the art, it is desirable to increase energy separation
between low and high kVp operation in order to increase the
contrast-to-noise ratio. However, it is not feasible to simply
decrease the low kVp or increase the high kVp in order to increase
energy separation therebetween. Lowering the low kVp may have
limited signal-to-noise and cause other limitations in image
reconstruction. Increasing the high kVp may cause system
instability and spit activity and may cause other limitations in
system operation.
Therefore, it would be desirable to have a system and method of
increasing energy separation in dual energy CT.
BRIEF DESCRIPTION
The present invention is directed to a system and method for
providing increased energy separation in dual energy CT.
According to an aspect of the present invention, an imaging system
includes an x-ray source that emits a beam of x-rays toward an
object to be imaged, a detector that receives the x-rays attenuated
by the object, a spectral notch filter positioned between the x-ray
source and the object, a data acquisition system (DAS) operably
connected to the detector, and a computer operably connected to the
DAS and programmed to acquire a first image dataset at a first kVp,
acquire a second image dataset at a second kVp that is greater than
the first kVp, and generate an image of the object using the first
image dataset and the second image dataset.
According to another aspect of the present invention, a method of
dual energy CT imaging includes selecting a low kVp potential and a
high kVp potential for dual energy imaging, selecting a k-edge
filter based on the low kVp potential and the high kVp potential
and based on a k-edge of a material in the k-edge filter,
positioning the k-edge filter between a source and an object to be
imaged, and acquiring imaging data with the source energized to the
first kVp potential and with the source energized to the second kVp
potential.
According to yet another aspect of the present invention, a method
of dual energy CT imaging includes passing low kVp x-rays through a
k-edge notch filter to generate a first x-ray spectrum, acquiring a
first set of imaging data of an object using the first x-ray
spectrum, passing high kVp x-rays through the k-edge notch filter
to generate a second x-ray spectrum, acquiring a second set of
imaging data of the object using the second x-ray spectrum, and
generating an image using the first set of imaging data and the
second set of imaging data.
Various other features and advantages of the present invention will
be made apparent from the following detailed description and the
drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
The drawings illustrate one preferred embodiment presently
contemplated for carrying out the invention.
In the drawings:
FIG. 1 is a pictorial view of a CT imaging system.
FIG. 2 is a block schematic diagram of the system illustrated in
FIG. 1.
FIG. 3 is a perspective view of one embodiment of a CT system
detector array.
FIG. 4 is a perspective view of one embodiment of a CT
detector.
FIG. 5 is an illustration of energy spectra that includes a low
energy spectrum and a high energy spectrum.
FIG. 6 is an illustration of a bowtie filter with k-edge material
according to an embodiment of the invention.
FIG. 7 is an illustration of a multi-material k-edge filter,
according to embodiments of the invention.
FIG. 8 is a pictorial view of a CT system for use with a
non-invasive package inspection system.
DETAILED DESCRIPTION
Imaging devices comprise x-ray systems, magnetic resonance (MR)
systems, ultrasound systems, computed tomography (CT) systems,
positron emission tomography (PET) systems, ultrasound, nuclear
medicine, and other types of imaging systems. Applications of x-ray
sources comprise imaging, medical, security, and industrial
inspection applications. It will be appreciated by those skilled in
the art that an implementation is applicable for use with
single-slice or other multi-slice configurations. Moreover, an
implementation is employable for the detection and conversion of
x-rays. However, one skilled in the art will further appreciate
that an implementation is employable for the detection and
conversion of other high frequency electromagnetic energy. An
implementation is employable with a "third generation" CT scanner
and/or other CT systems.
The operating environment of the present invention is described
with respect to a sixty-four-slice computed tomography (CT) system.
However, it will be appreciated by those skilled in the art that
the invention is equally applicable for use with other multi-slice
configurations. Moreover, the invention will be described with
respect to the detection and conversion of x-rays. However, one
skilled in the art will further appreciate that the invention is
equally applicable for the detection and conversion of other high
frequency electromagnetic energy. The invention will be described
with respect to a "third generation" CT scanner, but is equally
applicable with other CT systems.
Referring to FIG. 1, a computed tomography (CT) imaging system 10
is shown as including a gantry 12 representative of a "third
generation" CT scanner. Gantry 12 has an x-ray source 14 that
projects a beam of x-rays 16 through a bowtie filter 13 and toward
a detector assembly or collimator 18 on the opposite side of the
gantry 12. Referring now to FIG. 2, detector assembly 18 is formed
by a plurality of detectors 20 and data acquisition systems (DAS)
32. The plurality of detectors 20 sense the projected x-rays that
pass through a medical patient 22, and DAS 32 converts the data to
digital signals for subsequent processing. Each detector 20
produces an analog electrical signal that represents the intensity
of an impinging x-ray beam and hence the attenuated beam as it
passes through the patient 22. During a scan to acquire x-ray
projection data, gantry 12 and the components mounted thereon
rotate about a center of rotation 24.
Rotation of gantry 12 and the operation of x-ray source 14 are
governed by a control mechanism 26 of CT system 10. Control
mechanism 26 includes an x-ray controller 28 that provides power
and timing signals to an x-ray source 14 and a gantry motor
controller 30 that controls the rotational speed and position of
gantry 12. An image reconstructor 34 receives sampled and digitized
x-ray data from DAS 32 and performs high speed reconstruction. The
reconstructed image is applied as an input to a computer 36 which
stores the image in a mass storage device 38.
Computer 36 also receives commands and scanning parameters from an
operator via console 40 that has some form of operator interface,
such as a keyboard, mouse, voice activated controller, or any other
suitable input apparatus. An associated display 42 allows the
operator to observe the reconstructed image and other data from
computer 36. The operator supplied commands and parameters are used
by computer 36 to provide control signals and information to DAS
32, x-ray controller 28 and gantry motor controller 30. In
addition, computer 36 operates a table motor controller 44 which
controls a motorized table 46 to position patient 22 and gantry 12.
Particularly, table 46 moves patients 22 through a gantry opening
48 of FIG. 1 in whole or in part.
As shown in FIG. 3, detector assembly 18 includes rails 17 having
collimating blades or plates 19 placed therebetween. Plates 19 are
positioned to collimate x-rays 16 before such beams impinge upon,
for instance, detector 20 of FIG. 4 positioned on detector assembly
18. In one embodiment, detector assembly 18 includes 57 detectors
20, each detector 20 having an array size of 64.times.16 of pixel
elements 50. As a result, detector assembly 18 has 64 rows and 912
columns (16.times.57 detectors) which allows 64 simultaneous slices
of data to be collected with each rotation of gantry 12.
Referring to FIG. 4, detector 20 includes DAS 32, with each
detector 20 including a number of detector elements 50 arranged in
pack 51. Detectors 20 include pins 52 positioned within pack 51
relative to detector elements 50. Pack 51 is positioned on a
backlit diode array 53 having a plurality of diodes 59. Backlit
diode array 53 is in turn positioned on multi-layer substrate 54.
Spacers 55 are positioned on multi-layer substrate 54. Detector
elements 50 are optically coupled to backlit diode array 53, and
backlit diode array 53 is in turn electrically coupled to
multi-layer substrate 54. Flex circuits 56 are attached to face 57
of multi-layer substrate 54 and to DAS 32. Detectors 20 are
positioned within detector assembly 18 by use of pins 52.
In the operation of one embodiment, x-rays impinging within
detector elements 50 generate photons which traverse pack 51,
thereby generating an analog signal which is detected on a diode
within backlit diode array 53. The analog signal generated is
carried through multi-layer substrate 54, through flex circuits 56,
to DAS 32 wherein the analog signal is converted to a digital
signal.
Referring back to FIGS. 1 and 2, a discussion is now presented in
connection with a decomposition algorithm. An image or slice is
computed which may incorporate, in certain modes, less or more than
360 degrees of projection data to formulate an image. The image may
be collimated to desired dimensions using tungsten blades in front
of the x-ray source and different detector apertures. A collimator
typically defines the size and shape of the beam of x-rays 16 that
emerges from the x-ray source 14, and a bowtie filter 13 may be
included in the system 10 to further control the dose to the
patient 22. A typical bowtie filter attenuates the beam of x-rays
16 to accommodate the body part being imaged, such as head or
torso, such that, in general, less attenuation is provided for
x-rays passing through or near an isocenter of the patient 22. The
bowtie filter shapes the x-ray intensity during imaging in
accordance with the region-of-interest (ROI), field of view (FOV),
and/or target region of the patient 22 being imaged.
As the x-ray source 14 and the detector array 18 rotate, the
detector array 18 collects data of the attenuated x-ray beams. The
data collected by the detector array 18 undergoes pre-processing
and calibration to condition the data to represent the line
integrals of the attenuation coefficients of the scanned object or
the patient 22. The processed data are commonly called
projections.
In dual or multi-energy imaging, two or more sets of projection
data are typically obtained for the imaged object at different tube
peak kilovoltage (kVp) levels, which change the peak and spectrum
of energy of the incident photons comprising the emitted x-ray
beams or, alternatively, at a single tube peak kilovoltage (kVp)
level or spectrum with an energy resolving detector of the detector
array 18. The acquired sets of projection data may be used for
basis material decomposition (BMD). During BMD, the measured
projections are converted to a set of density line-integral
projections. The density line-integral projections may be
reconstructed to form a density map or image of each respective
basis material, such as bone, soft tissue, and/or contrast agent
maps (such as water and iodine). The density maps or images may be,
in turn, associated to form a volume rendering of the basis
material, for example, bone, soft tissue, and/or contrast agent, in
the imaged volume.
Once reconstructed, the basis material image produced by the CT
system 10 reveals internal features of the patient 22, expressed in
the densities of the two basis materials. The density image may be
displayed to show these features. In traditional approaches to
diagnosis of medical conditions, such as disease states, and more
generally of medical events, a radiologist or physician would
consider a hard copy or display of the density image to discern
characteristic features of interest. Such features might include
lesions, sizes and shapes of particular anatomies or organs, and
other features that would be discernable in the image based upon
the skill and knowledge of the individual practitioner.
In addition to a CT number or Hounsfield value, an energy selective
CT system can provide additional information related to a
material's atomic number and density. This information may be
particularly useful for a number of medical clinical applications,
where the CT number of different materials may be similar but the
atomic number may be quite different. For example, calcified plaque
and iodine-contrast enhanced blood may be located together in
coronary arteries or other vessels. As will be appreciated by those
skilled in the art, calcified plaque and iodine-contrast enhanced
blood are known to have distinctly different atomic numbers, but at
certain densities these two materials are indistinguishable by CT
number alone.
A decomposition algorithm is employable to generate atomic number
and density information from energy sensitive x-ray measurements.
Multiple energy techniques comprise dual energy, photon counting
energy discrimination, dual layered scintillation and/or one or
more other techniques designed to measure x-ray attenuation in two
or more distinct energy ranges. As an example, a compound or
mixture of materials measured with a multiple energy technique may
be represented as a hypothetical material (or combination of
materials) having the same x-ray energy attenuation
characteristics. This hypothetical material can be assigned an
effective atomic number Z. Unlike the atomic number of an element,
effective atomic number of a compound is defined by the x-ray
attenuation characteristics, and it need not be an integer. This
effective Z representation property stems from a well-known fact
that x-ray attenuation in the energy range useful for diagnostic
x-ray imaging is strongly related to the electron density of
compounds, which is also related to the atomic number of
materials.
Thus, dual-energy CT with fast kVp switching is an attractive way
of achieving near simultaneous and near co-registered projection
samples of two energies. However, because of the fast switching,
there is little opportunity to change filtration between samples or
otherwise increase energy separation between the low and high kVp
energies. Thus, according to an embodiment of the invention, a
single filter with an energy notch or k-edge in an overlapped
region of the low and high kVp energies may be employed to increase
energy separation therebetween.
Referring to FIG. 5, an illustration 100 of energy spectra includes
a low energy spectrum 102 having a first peak energy 104 and a high
energy spectrum 106 having a second peak energy 108. Low energy
spectrum 102 includes a first mean keV 110 and high energy spectrum
106 includes a second mean keV 112. An amount of energy separation
114 is illustrated between first mean keV 110 and second mean keV
112. As known in the art, each mean keV 110, 112 represents an
energy or keV that approximately splits an amount of integrated
area for each respective spectrum. Thus, for low energy spectrum
102, first mean keV 110 represents an energy level where an
integrated energy below mean 116 is approximately equal to an
integrated energy 118 above mean 116. Similarly, high energy
spectrum 106 includes integrated energies (not marked) below and
above second mean keV 112.
Spectra 102, 106 represent energy spectra emitted from an x-ray
tube at respective peak energies 104, 108. In one example, a
typical representation of respective energies is for equivalent
patient filtration for an amount of water thickness. Thus, in this
example that includes an equivalent patient filtration of 20 cm
water, for a peak low kVp of 80 keV and for a peak high kVp of 140
keV, mean energies are approximately 55 keV and 76 keV,
respectively. This results in an approximate energy separation of
21 keV (76 keV minus 55 keV) between low and high spectra.
However, when placing a k-edge material between an x-ray source and
detector according to the invention, it is possible to increase
energy separation between mean low kVp and mean high kVp. A k-edge
indicates a sudden increase in the attenuation coefficient of
photons occurring at a photon energy just above the binding energy
of the K shell electron of the atoms interacting with the photons.
The sudden increase in attenuation is due to photoelectric
absorption of the photons. For this interaction to occur, the
photons have more energy than the binding energy of the K shell
electrons. A photon having an energy just above the binding energy
of the electron is therefore more likely to be absorbed than a
photon having an energy just below this binding energy. A general
term for the phenomenon is absorption edge.
Because of this sudden jump in attenuation, it is possible to
increase separation of mean energies of low and high kVp spectra,
according to the invention. In one example, for 20 cm of water and
0.5 mm Hf (k-edge of approximately 65.4 keV), mean energies of low
and high are respectively, approximately 58 keV and 86 keV,
resulting in a separation of approximately 28 keV--which is an
increase from 21 keV as illustrated above. Referring back to FIGS.
1 and 2, k-edge material 15 may be positioned between x-ray source
14 and detector assembly 18, and more particularly between x-ray
source 14 and patient 22. As such, with placement of an attenuating
material having a k-edge that falls between mean energies of the
low kVp and high kVp spectra it is possible to increase separation
therebetween, according to the invention.
Thus, in general and according to the invention, it is possible to
increase energy separation between low and high kVp spectra by
selecting a k-edge notch filter having a k-edge that falls between
the mean energies of the low and high kVp spectra. Typically, such
a filter may have a thickness of approximately 1 mm. However, it is
to be understood that the thickness is dependent on specific
desired imaging characteristics including but not limited to low
and high kVp spectra, mA, patient characteristics, anatomy, and the
like.
In an example, low and high kVp spectra are respectively 80 keV and
140 keV. And, in one example the low kVp potential and the high kVp
potential are each for a period less than one millisecond. However,
it is to be understood that any low and high kVp spectra may be
selected for dual or multi-energy imaging, according to the
invention. It is also to be understood that one millisecond
duration at low and high kVp potentials is an example, and that any
length period may be implemented, depending on imaging application,
according to the invention. Further, although Hf is given above as
an example k-edge material, according to the invention any material
having a k-edge between mean low kVp 110 and mean high kVp 112 may
suffice. Thus, for dual energy imaging, typical desired k-edge
materials may range between approximately 30 keV and 80 keV.
Although bowtie filter 13 and k-edge filter 15 are illustrated as
separate elements FIGS. 1 and 2, it is possible to combine both
into a single apparatus that includes both bowtie filtration and
k-edge filtration. Referring now to FIG. 6, a bowtie filter is
illustrated according to an embodiment of the invention. Typically,
a bowtie filter may include multiple bowties that may be accessed
by selectively placing the bowtie filter at a preferred axial
location. FIG. 6 is an illustration of one example of a bowtie
filter unit 200 having two sizes of bowtie 202, 204 therein. Each
bowtie 202, 204 is positioned along an axis 206 of bowtie filter
unit 200. Thus, when in operation, bowtie filter unit 200 may be
selectively placed axially, based on an anatomy that is to be
imaged, or based on a patient that is to be imaged. As such, in one
example for a relatively small body, bowtie filter 202 may be
selected, while for a relatively large body, bowtie filter 204 may
be selected. And, bowtie filter unit 200 may not be limited to two
sizes of bowtie 202, 204, but may include many bowties that are
positionable along axis 206. Bowtie filter 200 may include a k-edge
material that can serve a dual purpose of providing bowtie beam
shaping as well as k-edge filtration. Thus, bowtie filter 200 may
include a k-edge material in one or both bowties 202, 204, as
illustrated in phantom as an example as k-edge material 208.
It is possible to enhance filtration, selectability, separation of
spectra, and more controlled shaping of the energy spectrum by
combining two or more k-edge filters. Thus, according to
embodiments of the invention, two or more k-edge materials may be
included in a k-edge filter. FIG. 7 is thus an illustration of a
multi-material k-edge filter, according to an embodiment of the
invention. FIG. 7 is an illustration of x-rays 300 that emanate
from a focal spot or point 302, which may be for instance a focal
spot or point within an x-ray source such as x-ray tube 14 of FIGS.
1 and 2. X-rays 300 pass through a multi-material k-edge filter
304, through a patient or object (not shown), and toward a detector
array or assembly 306 (which may be, in an example, detector
assembly 18 of FIGS. 1 and 2).
As illustrated, multi-material k-edge filter 304 includes a first
material 308 and a second material 310. In embodiments of the
invention, first and second materials 308, 310 are k-edge materials
that, in combination, enable a selective and controlled shaping of
the energy spectrum, which can lead to distinct notch filtration
when compared to a single k-edge material. As such, a combination
of k-edge materials may be selected in order to specifically affect
a level of filtration at a specific energy, while leaving portions
of the spectrum outside this specifically affected area intact.
And, although two materials 308, 310 are illustrated, it is to be
understood that more than two materials may be included, limited
only by a combined and desired total attenuation and space
available for placement of multi-material k-edge filter 304. It is
to be understood that the use of multi-material k-edge filter 304
may be in combination with a separate conventional bowtie filter,
or multi-material k-edge filter 304 may be combined with a bowtie
filter to provide both k-edge filtration and bowtie beam shaping in
a single unit. As stated above with respect to a single k-edge
material filter, combinations of materials may be selected, each
having for instance a k-edge that falls between approximately 30
keV and 80 keV.
In one example, Hf (hafnium) and W (tungsten) may be combined to
enable an improved optimization over Hf alone. This combination of
k-edge materials allows selective choice of a region of the
spectrum that is desired to be affected to tune attenuation in that
region of the spectrum. Further, Hf and W are used as an example of
a combination of materials. However, depending on the energy range
that is desired to be affected, different materials can be selected
based on their k-edge, density, and the like, which can be combined
to create the effect at selected energy ranges in the spectrum.
Referring now to FIG. 8, package/baggage inspection system 500
includes a rotatable gantry 502 having an opening 504 therein
through which packages or pieces of baggage may pass. The rotatable
gantry 502 houses an x-ray and/or high frequency electromagnetic
energy source 506 as well as a detector assembly 508 having
scintillator arrays comprised of scintillator cells. A conveyor
system 510 is also provided and includes a conveyor belt 512
supported by structure 514 to automatically and continuously pass
packages or baggage pieces 516 through opening 504 to be scanned.
Objects 516 are fed through opening 504 by conveyor belt 512,
imaging data is then acquired, and the conveyor belt 512 removes
the packages 516 from opening 504 in a controlled and continuous
manner. As a result, postal inspectors, baggage handlers, and other
security personnel may non-invasively inspect the contents of
packages 516 for explosives, knives, guns, contraband, etc. An
exemplary implementation can aid in the development of automatic
inspection techniques, such as explosive detection in luggage.
A technical contribution for the disclosed method and apparatus is
that is provides for a computer implemented system and method of
basis material decomposition having an increased separation of mean
energies between low and high kVp projections.
One skilled in the art will appreciate that embodiments of the
invention may be interfaced to and controlled by a computer
readable storage medium having stored thereon a computer program.
The computer readable storage medium includes a plurality of
components such as one or more of electronic components, hardware
components, and/or computer software components. These components
may include one or more computer readable storage media that
generally stores instructions such as software, firmware and/or
assembly language for performing one or more portions of one or
more implementations or embodiments of a sequence. These computer
readable storage media are generally non-transitory and/or
tangible. Examples of such a computer readable storage medium
include a recordable data storage medium of a computer and/or
storage device. The computer readable storage media may employ, for
example, one or more of a magnetic, electrical, optical,
biological, and/or atomic data storage medium. Further, such media
may take the form of, for example, floppy disks, magnetic tapes,
CD-ROMs, DVD-ROMs, hard disk drives, and/or electronic memory.
Other forms of non-transitory and/or tangible computer readable
storage media not list may be employed with embodiments of the
invention.
A number of such components can be combined or divided in an
implementation of a system. Further, such components may include a
set and/or series of computer instructions written in or
implemented with any of a number of programming languages, as will
be appreciated by those skilled in the art. In addition, other
forms of computer readable media such as a carrier wave may be
employed to embody a computer data signal representing a sequence
of instructions that when executed by one or more computers causes
the one or more computers to perform one or more portions of one or
more implementations or embodiments of a sequence.
Therefore, according to an embodiment of the invention, an imaging
system includes an x-ray source that emits a beam of x-rays toward
an object to be imaged, a detector that receives the x-rays
attenuated by the object, a spectral notch filter positioned
between the x-ray source and the object, a data acquisition system
(DAS) operably connected to the detector, and a computer operably
connected to the DAS and programmed to acquire a first image
dataset at a first kVp, acquire a second image dataset at a second
kVp that is greater than the first kVp, and generate an image of
the object using the first image dataset and the second image
dataset.
According to another embodiment of the invention, a method of dual
energy CT imaging includes selecting a low kVp potential and a high
kVp potential for dual energy imaging, selecting a k-edge filter
based on the low kVp potential and the high kVp potential and based
on a k-edge of a material in the k-edge filter, positioning the
k-edge filter between a source and an object to be imaged, and
acquiring imaging data with the source energized to the first kVp
potential and with the source energized to the second kVp
potential.
According to yet another embodiment of the invention, a method of
dual energy CT imaging includes passing low kVp x-rays through a
k-edge notch filter to generate a first x-ray spectrum, acquiring a
first set of imaging data of an object using the first x-ray
spectrum, passing high kVp x-rays through the k-edge notch filter
to generate a second x-ray spectrum, acquiring a second set of
imaging data of the object using the second x-ray spectrum, and
generating an image using the first set of imaging data and the
second set of imaging data.
The present invention has been described in terms of the preferred
embodiment, and it is recognized that equivalents, alternatives,
and modifications, aside from those expressly stated, are possible
and within the scope of the appending claims.
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